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This article was published as part of the
Nanomedicine themed issue
Guest editors Frank Caruso, Taeghwan Hyeon and Vincent Rotello
Please take a look at the issue 7 2012 table of contents to
access other reviews in this themed issue
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TUTORIAL REVIEW
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Polysaccharides as building blocks for nanotherapeuticsw
Shoshy Mizrahyab and Dan Peer*ab
Received 5th September 2011
DOI: 10.1039/c1cs15239d
The use of polysaccharides as building blocks in the development of nano-sized drug delivery
systems is rapidly growing. This can be attributed to the outstanding virtues of polysaccharides
such as biocompatibility, biodegradability, low toxicity and low cost. In addition, the variety
of physicochemical properties and the ease of chemical modifications enable the preparation
of a wide array of nanoparticles. This tutorial review describes the properties of common
polysaccharides, the main mechanisms for polysaccharide based-nanoparticles preparation, and
provides examples from the conceptual design towards pre-clinical and clinical applications.
Laboratory of Nanomedicine, Dept. of Cell Research and
Immunology, Georg S. Wise Faculty of Life Science,
Tel Aviv University, Tel Aviv, 69978, Israel.
E-mail: [email protected]; Fax: +972-3-6405926;
Tel: +972-3-6407925
b
Center for Nanoscience and Nanotechnology, Tel Aviv University,
Tel Aviv, Israel
w Part of the nanomedicine themed issue.
mechanisms, improved biodistribution and pharmacokinetics and
enhanced intracellular penetration.2
The size, geometrical shape and surface characteristics of a
NP are crucial for the control of its biodistribution in vivo.1
The small size, which enables NPs to pass through the smallest
capillaries, also promotes passive tumor targeting due to the
enhanced permeability and retention (EPR) effect of the tumor
vasculature (Fig. 1). The passive targeting is achieved by
extravasation of NPs through increased permeability of the
newly formed tumor blood vessels, caused by rapid and defective angiogenesis.2 In addition, the impaired tumor lymphatic
drainage leads to accumulation of NPs in tumor surroundings
in a size dependent manner.2 Cellular targeting can be achieved
by binding a targeting agent (ligand) onto the NPs surface, this
will allow the particles to actively bind to specific cells following
extravasation (Fig. 1).2
Shoshy Mizrahy received her
BSc in Biotechnology and
Food Engineering from the
Technion, Israel Institute of
Technology in 2002. From
2002–2003 she worked as a
research assistant at BioTechnology General (Israel)
Ltd. She then obtained her
MSc in Pathology from Tel
Aviv University School of
Medicine
(2006).
From
2006–2009 she worked as a
researcher in Procognia Ltd.
focusing on glycobiology and
Shoshy Mizrahy
glycochemistry. From 2009
she is a PhD graduate student in the Laboratory of Nanomedicine at Tel Aviv University focusing on novel nano-sized
delivery systems made from polysaccharides for the delivery of
therapeutic payloads to hematological malignancies.
Dan Peer received his PhD in
biophysics and biochemistry
from Tel Aviv University.
From 2005–2008 he did his
postdoctoral studies at Harvard
Medical School. During his
time at Harvard he developed
several novel strategies to
deliver siRNAs and other therapeutic payloads to immune cells
using targeted nanoparticles.
In 2008 he established the
Laboratory of Nanomedicine
at Tel Aviv University. His
current research includes
Dan Peer
developing novel strategies for
targeted nanomedicines based on polysaccharides, probing and
manipulating the immune system with nanomaterials, harnessing
RNAi as a tool for drug discovery and for therapeutic applications, and developing tools to study immuno-nanotoxicity.
1. Introduction
Over the past two decades nanoparticles (NPs)-based therapeutics
have been introduced for the treatment of cancer, diabetes, allergy,
neurodegenerative diseases, infections and inflammation.1,2 The
growing interest in NPs derives from the outstanding advantages
they offer, which include protection of the drug from premature
degradation, the ability to deliver poorly-water soluble drugs alone
or in combination with soluble drugs, controlled drug release
a
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Fig. 1 Passive and active tumor targeting. Passive tissue targeting is achieved by extravasation of nanoparticles through increased permeability of
the tumor vasculature and ineffective lymphatic drainage (EPR effect). Active cellular targeting (inset) can be achieved by functionalizing the
surface of nanoparticles with ligands that promote cell-specific recognition and binding. The nanoparticles can (i) release their contents in close
proximity to the target cells; (ii) attach to the membrane of the cell and act as an extracellular sustained-release drug depot; or (iii) internalize into
the cell. Reprinted with permission from ref. 2. Copyright 2007 Nat. Nanotechnol.
The effect of spherical, rigid NPs size on circulation time has
been studied extensively and size ranges between 100–200 nm
have been determined optimal for long circulation.1 These NPs
are large enough to avoid renal clearance and liver uptake and
in the appropriate size to evade uptake by the mononuclear
phagocyte system (MPS).1,3 The particle size, shape and the
presence of surface ligands also play a crucial role in intracellular delivery since it affects the mechanism of cellular internalization.1,2 The specific endocytotic mechanism is significant
since it determines the intracellular path and microenvironment
encountered by the NPs. For example, in clathrin-mediated
endocytosis, which is characteristic for specific particle sizes
and also for certain ligands, NPs should be capable of endosomal escape in order to avoid the harsh lysosomal environment.1
The effects of particle geometry on cell internalization and
biodistribution are just being revealed and yet a recent literature points out that the particles’ geometry is as crucial as
its size.1
The surface feature of the NP is another major factor in
determining the biodistribution. Poly(ethylene glycol) (PEG),
which has been initially used to promote long circulation of
therapeutic proteins, has been shown to elongate the circulation time of NPs. The PEG coating alters the hydrophobicity
of the NP’s surface protecting it from protein opsonisation—a
process that marks the NP for removal from the circulation by
specialized macrophages.2
When designing a NP as a therapeutic entity, the mechanism
for release of the therapeutic cargo should also be considered.
Several activated release methods have been utilized to break
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the bond between the drug and NP leading to particle degradation,
among which are enzymatic cleavage, pH or thermal instability.1
As the requirements from NPs as therapeutic carriers are
becoming clear so are the requirements from the materials
used for their preparation. These materials should be biocompatible and biodegradable, well characterized and easily functionalized.2 Polysaccharides successfully fulfill all of these
requirements and are therefore widely used as the building
blocks for the preparation of NPs as drug delivery vehicles.
This tutorial review details the properties of polysaccharides
used for the preparation of drug delivery vehicles, evaluates the
challenges and discusses the opportunities in this exciting field.
2. Polysaccharides
Polysaccharides are polymers of monosaccharides joined by glycosidic bonds. These highly abundant molecules are from various
origins including algal origin (e.g. alginate and carrageenan), plant
origin (e.g. cellulose, pectin and guar gum), microbial origin
(e.g. dextran and xanthan gum), and animal origin (e.g. CS,
hyaluronan, chondroitin and heparin).4 Naturally accruing
polysaccharides are diverse in their physiochemical properties;
there are multiple chemical structures (Fig. 2), the chemical
composition greatly varies and so are the molecular weight
(Mw) and ionic nature. This versatility also contributes to a
wide range of biological activities. From a pharmaceutical
standpoint, polysaccharides possess many favorable characteristics such as low toxicity, biocompatibility, stability, low
cost, hydrophilic nature and availability of reactive sites for
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Fig. 2 Chemical structure of polysaccharides. Chitosan, the N-deacetylated derivative of chitin is composed of b (1,4)-N-acetyl-D-glucosamine
and D-glucosamine. Hyaluronan is composed of alternating b-(1,4)-D-glucuronic acid and b-(1,3)-N-acetyl-D-glucosamine. Alginate is composed of
alternating blocks of b-(1,4)-D-mannuronic acid and a-(1,4)-L-guluronic acid. Chondroitin sulfate is composed of alternating b-(1,3)-N-acetylD-galactosamine and b-(1,4)-glucuronic acid. Amylose is composed of a-(1,4)-D-glucose units. Heparin is composed of a or b-(1,4) linked uronic
acid (90% a-L-iduronic acid, 10% b-D-glucuronic acid) and a-D-glucosamine residues.
chemical modification.4,5 Chemical functionalization using
mainly the free carboxyl and hydroxyl groups distributed
along the polysaccharides backbone has been used to create
derivatives with determined/tailored properties.6 The solubility,
hydrophobicity, physicochemical and biological characteristics
of polysaccharides have been modified. This was performed by
using techniques such as oxidation, sulfation, esterification,
amidation, or grafting methods.6
Another advantage of polysaccharides is bioadhesion, especially
for mucosal surfaces, which has been used for targeting specific
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organs or cells and prolonging the drug residence time. All of
these qualities have led to the growing use of polysaccharides in
drug delivery systems.
The properties of common polysaccharides used for the
preparation of drug delivery systems are detailed below.
2.1 Chitosan
Chitosan (CS) is a linear polysaccharide composed of b-(1,4)linked D-glucosamine and N-acetyl-D-glucosamine (Fig. 2).
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CS is obtained by deacetylation from chitin, a highly abundant
polysaccharide, which is the main component of the crustaceans exoskeleton.7 CS based delivery systems have been
described for nasal, ocular, oral, parenteral and transdermal
drug delivery. Among the many advantages of CS are its low
cytotoxicity and biocompatibility.7 As such, CS is approved
for dietary applications in Japan, Italy and Finland and it
has been approved by the FDA for use in wound dressings.8
Interestingly, CS and its derivatives were found to be toxic to
several bacteria, fungi and parasites,8 another welcome trait
when aiming for delivery of pharmaceuticals. In addition, CS
is positively charged and therefore can interact with negatively
charged molecules such as negatively charged polysaccharides,
polyanions, nucleic acid and negatively charged proteins.9 Its
positive charge also facilitates adherence to mucosal surfaces,
which are mostly negatively charged.7 In spite of these advantages, CS is inherently insoluble in aqueous solutions above
pH 6.5.10 High degree of deacetylation, low molecular weight
and chemical modification can facilitate water solubility of CS.
These factors also affect particle properties such as size, surface
charge, drug entrapment efficiency, and stability.7
Comparison of CS and its derivatives to synthetic cationic
polymers such as polyethylenimine (PEI) reveals significantly
lower toxicity. However, upon increasing CS charge density in
order to improve cell uptake, its toxicity increased8 as can be
expected when taking into consideration the charge related
adverse effects of positively charged synthetic lipids and polymers.
Nevertheless, this is not necessarily problematic since a significant improvement in cell uptake can be achieved simply by
grafting ligands that mediate endocytosis onto CS or any other
polysaccharide. CS derivatives that do not increase charge
density and all low Mw CS and its derivatives (o10 kDa) are
not appreciably toxic.8 Yet, every newly formed derivative
should be examined for toxicity and used in the purest form.8
This justifies the selection of CS (and its derivatives) that
provides health benefits over current formulations as a safe
material in drug delivery.8
2.2
Alginate
Alginate is a linear anionic polysaccharide composed of alternating blocks of 1,4-linked b-D-mannuronic acid (M) and
a-l-guluronic acid (G) residues (Fig. 2). The monomer composition of alginate is variable and can consist of homopolymeric
blocks and alternating M and G residues.11 The composition,
sequence, and molecular weight determine the physical properties of alginate.11 Alginates are extracted mainly from brown
algae and acetylated forms of alginate can be isolated from the
bacteria Pseudomonas and Azotobacter.11
As a polymer used in drug delivery, alginate possesses several
attractive properties: it is biocompatible, non-toxic, water soluble
and highly mucoadhesive.11
2.3
Hyaluronan
Hyaluronan (HA) is a linear high Mw glycosaminoglycan (GAG)
composed of alternating disaccharide units of D-glucuronic acid
and N-acetyl-D-glucosamine with b-(1, 4) interglycosidic linkage12
(Fig. 2). Hyaluronan holds remarkable hydrodynamic properties
especially regarding its viscosity and ability to retain water.12
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It was previously regarded as important for joint lubrication
and organ structural stability.12 However, HA was found to be
essential for proper cell growth, embryonic development,
healing processes, inflammation, and tumor development.12,13
As opposed to other GAGs, HA is not sulfated, not linked to
a protein12 and is naturally produced by bacteria as a capsule.
Commercially available HA is either produced through bacterial
fermentation of Streptococcus species or extracted from rooster
combs, umbilical cords, synovial fluids or vitreous humour.
There are several advantages of HA, which make it suitable
for drug delivery: it is water soluble, biodegradable, biocompatible, non-toxic, non-immunogenic and can be easily chemically
modified.13 In addition, it is the major ligand for CD44 and
CD168 (also known as Receptor for Hyaluronan Mediated
Motility, RHAMM) and therefore is suitable for targeting
CD44 and RHAMM-expressing cells.13 CD44 and CD168 are
overexpressed by various tumors, for example, squamous cell
carcinoma, ovarian, colon, stomach, glioma, and many types of
leukemia, lymphoma and multiple myeloma, which make the
use of HA as a targeting agent even more attractive.
2.4 Dextran
Dextran is a high Mw polysaccharide composed of a-(1,6)linked glucan with side chains attached to the 3-positions of
the backbone glucose units (Fig. 2). Dextran is obtained from
bacterial cultures of lactic acid bacteria such as Leuconostoc
mesenteroides NRRL B-512.4 Dextran is water-soluble and
also soluble in a wide range of solvents.
The presence of dextran degrading enzymes derived from
anaerobic gram negative intestinal bacteria is a motivating
factor for the development of dextran based NPs for colon
targeted drug delivery.4 Dextran bears several advantages as
a polymer for drug delivery: it is highly water soluble, biocompatible, biodegradable, lacks nonspecific cell binding and
resistant to protein adsorption.5 In addition, dextran is easily
functionalized via its reactive hydroxyl chemistries.5
2.5 Cyclodextrins
Cyclodextrins (CDs) are natural cyclic oligomers of a-(1,4)
linked-glucopyranosyl that are produced from starch by enzymatic conversion (Fig. 2). There are three main members of the
CD family, composed of six, seven and eight glucose units and
known as a-, b- and g-CD, respectively. CDs have a hydrophilic
exterior and a hydrophobic cavity that enables them to act as
hosts to hydrophobic molecules.14,15 This unique ability to form
inclusion complexes has been utilized by the pharmaceutical
industry to improve bioavailability of poorly soluble or biodegradable drugs and to enhance permeability of biological
membranes.14 In addition to stabilization of small drug molecules,
CDs have been shown to increase the stability of oligonucleotides
against endonucleases and even to modulate undesirable side
effects such as immune stimulation.14 The mechanism for this
is not completely clear but may be attributed to the fact that
CDs have the ability to stabilize biomolecules in physiological
media by shielding them from non-specific interactions.14 This
shielding ability may explain the protection against nucleases
and interactions with specific proteins that may lead to immune
stimulation.16 In addition, CDs have been shown to have
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membrane permeation capabilities.14 This can attribute to
lowering the amount of time the oligonucleotides are found
in the circulation and exposed to nucleases.
CDs are biocompatible, do not elicit an immune response,
and have low toxicities in animals and humans.15 Therefore,
they are used in pharmaceutical applications for numerous
purposes, including improving the bioavailability of drugs.15
CDs based therapeutics have been reviewed elsewhere.14,15
2.6
Arabinogalactan
Arabinogalactan is a long, highly branched natural polysaccharide composed mostly of galactose and arabinose.
Arabinogalactan is extracted mainly from the Larix tree and
is available at 99.9% purity with reproducible molecular
weight (Mw) and physicochemical properties.17 The unusual
water solubility (70% w/w in water), biocompatibility, biodegradability and ease of drug conjugation in an aqueous
medium makes arabinogalactan attractive as a potential drug
carrier.17
2.7
Pullulan
Pullulan is neutral, homopolysaccharide consisting of a–(1,6)linked maltotriose residues (Fig. 2). It is produced from starch
primarily by strains of the fungus Aureobasidium pullulans.18
The unique linkage pattern of pullulan’s structure contributes
to exceptional physiochemical properties such as adhesiveness,
water solubility and relatively low viscosity upon dissolving in
water. Therefore, pullulan and its derivatives have been used
industrially in foods and pharmaceuticals.18
2.8
Heparin
Heparin is a linear glycosaminoglycan (GAG) composed of
repeating disaccharide units of 1,4-linked uronic acid (D-glucuronic
(GlcA) or l-iduronic acid (IdoA)) and D-glucosamine (GlcN)
(Fig. 2). The uronic acid usually comprises 90% l-idopyranosyluronic acid (l-iduronic acid, IdoA) and 10% D-glucopyranosyluronic acid (D-glucuronic acid, GlcA). In addition, there are
structural variations at the disaccharide level.19 Due to high
content of sulfo and carboxyl groups, heparin has the highest
negative charge density of any known biological molecule.19
The Mw of heparin varies between 5–40 kDa and it is extracted
mainly from mucosal tissues of porcine and bovine.19 Clinically,
heparin has been used as an anticoagulant since the 1930s.19
Heparin is produced exclusively by mast cells (as opposed to the
structurally related GAG heparan sulfate). Beyond its anticoagulant activity, heparin has been shown to have antiviral
activity, regulate angiogenesis and inhibit complement activation.20
3. The main mechanisms of nanoparticle
preparation from polysaccharides
3.1
Cross-linking
In crosslinked NPs, the polymeric chains are interconnected
by crosslinkers, leading to the formation of a 3D network
(Fig. 3A and B).21 The main factor, which determines the
properties of a crosslinked NP such as drug release and mechanical strength, is the cross-linking density, which is determined
by the molar ratio between the crosslinker and the polymer
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Fig. 3 Common mechanisms for polysaccharide based nanoparticle
preparation. (A) Covalent cross-linking. (B) Ionic cross-linking.
(C) Polyelectrolyte complexation (PEC). (D) Self-assembly of hydrophobically modified polysaccharides. (E) Polysaccharide–drug conjugate.
repeating units.21 There are two kinds of crosslinked NPs
determined by the nature of the cross-linking agents:
covalently crosslinked NPs and ionically crosslinked NPs.
3.1.1 Covalent cross-linking. In a covalently cross-linked
NP, the network structure is permanent since irreversible
chemical links are formed unless biodegradable or stimuliresponsive crosslinkers are employed (Fig. 3A).21 The rigid
network allows absorption of water and bioactive compounds
without dissolution of the NP even when the pH drastically
changes.21 A covalently crosslinked NP can contain more than
one type of a polysaccharide. The covalent bonds are the main
interactions that form the 3D network although secondary
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interactions such as hydrogen bonds and hydrophobic interactions also exist.21 Covalent crosslinkers are molecules with
at least two reactive functional groups that allow the formation of bridges between the polymeric chains.22 The most
common covalent crosslinkers used with polysaccharides are
dialdehydes such as glutaraldehyde. However, dialdehydes are
highly toxic and therefore biocompatible alternatives have
been tested. For example, natural di- and tricarboxylic acids
have been used for intramolecular cross-linking of CS, which
was facilitated by the condensation agent N,N-(3-dimethylaminopropyl)-N-ethyl carbodiimide (EDC).9 Another alternative is Genipin, a natural biocompatible crosslinker isolated
from the fruits of Gardenia jasminoides Ellis.21
3.1.2 Ionic cross-linking. Ionic cross-linking represents a
simple alternative to covalent cross-linking for charged polysaccharides. This method enables the preparation of NPs by
the formation of reversible ionic cross-linking and since no
harsh preparation or toxic crosslinkers are used these NPs
are generally considered biocompatible.21 Charged polysaccharides can form ionic crosslinked NPs with oppositely
charged ions or small ionic molecules (Fig. 3B). For example,
the polyanion tripolyphosphate (TPP) has been widely used to
crosslink the CS, and divalent cations such as Ca2+ have been
used to crosslink alginate.7 The mechanism of NP formation is
based on electrostatic interactions between the polysaccharide
and oppositely charged ionic crosslinker. The ionic bonds that
form bridges between the polysaccharide chains are the main
interactions inside the network, although as with covalent
cross-linking, additional interactions such as hydrogen bonds
are also present.21 Several factors influence the cross-linking
reaction, most crucial are the size of the crosslinker and the
global charge of the crosslinker and the polysaccharide.21
Unlike covalently crosslinked NPs, ionic crosslinked particles
are generally pH sensitive, a welcome trait for drug delivery
purposes. However, this pH sensitivity can also contribute to
instability of the ionic crosslinked network.21
3.2
Polyelectrolyte structures
3.2.1 Polyelectrolyte complexes (PEC). Polyelectrolyte
complexes (PEC) are formed by direct electrostatic interactions
of oppositely charged polyelectrolytes in solution (Fig. 3C).
PEC represent another biocompatible option for drug delivery
since non-toxic covalent crosslinkers are used. These complexes
resemble ionic cross-linking since non-permanent networks are
formed that are more sensitive to changes in environmental
conditions.23 However, unlike ionic cross-linking, in which ions
or ionic molecules react with the polyelectrolyte, in PEC the
interaction is between the polyelectrolyte and larger molecules
with broad Mw range.21 The formation and stability of PEC
are determined mainly by the degree of interaction between the
polyelectrolytes.22 The latter is a factor of the charge density and
distribution of each of the oppositely charged polyelectrolyte.
The chemical environment is also crucial: the pH of the solution,
the ionic strength, the temperature, and the duration and mixing
order. Secondary factors are the Mw of the polyelectrolytes and
their flexibility.22,24 Ionic cross-linking can reinforce the formed
interaction. Positively charged polysaccharides, namely CS, can
form PEC with a variety of negatively charged polymers such as
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the polysaccharides alginate, dextran sulfate, chondroitin sulfate,
hyaluronan, carboxymethyl cellulose, carrageenan and heparin.
In addition, peptides such as poly-g-glutamic acid, nucleic acids
and synthetic polymers could also be used.21,24
3.2.2 Layer by layer assembly. Layer by layer (LbL)
assembly of polyelectrolyte NPs is a relatively new and therefore less common form of polyelectrolyte-based nano-sized
delivery systems. The LbL technique was originally used for
generating thin polyelectrolyte layers on solid surfaces and has
shown broad biomedical applications in biosensing, regenerative medicine, tissue engineering, and biomimetics research.25 This
technique, which is based on electrostatic interactions, employs
alternate adsorption of oppositely charged polyelectrolytes.26
The obtained film structure can be controlled to 1 nm precision with a range from 5 to 1000 nm and with a definite
knowledge of the molecular composition (see also Fig. 9C).26
The assembly procedure involves the incubation of charged
solid support in an oppositely charged polyelectrolyte solution.
This results in the adsorption of the first layer, which is
subsequently rinsed to remove excess free polyelectrolyte and
immersed in a solution containing a second polyelectrolyte to
create the second layer. The process is repeated until the desired
thickness is obtained.26 The polyelectrolyte adsorption at each
layer is performed to a level of saturation and therefore the
terminal charge is alternated after every subsequent layer
deposition.26 The demonstration of multilayer assembly on a
three-dimensional support presented this technique for applications in nano-sized delivery systems.25 The advantages for the
LbL approach are clear: many therapeutics and biomaterials
can be incorporated into LbL films non-covalently and under
physiological conditions, without compromising their biological
properties.25 In addition, since the technique can be used with
multiple components, several factors can be manipulated,
among them are the surface chemistry, dimensions of the thin
films, therapeutic moieties used and, thus a sophisticated nanosized drug delivery system can be created with a tailored drug
release mechanism.25 Furthermore, LbL NPs can be used for
the delivery of multiple active substances with the ability to
control the release of each entrapped substance.25 The possibility to reach sequential drug release from LbL film layers enables
the control of the order and timing of multiple drug release.25
3.3 Self-assembly
Upon grafting hydrophobic moieties onto a hydrophilic polysaccharide, an amphiphilic copolymer is created. In aqueous
solutions amphiphilic copolymers tend to self-assemble into
NPs in which the inner core is hydrophobic and the shell is
hydrophilic. The hydrophilic shell serves as a stabilizing interface between the hydrophobic core and the external aqueous
environment (Fig. 3D).27 This self-assembly process is via
hydrophobic interactions, mainly in order to minimize interfacial free energy.9,24 The formed NPs are characterized by
prolonged circulation and thermodynamic stability.28 In addition, since the core is hydrophobic, these particles have been
used for the delivery of hydrophobic drugs. Several properties
such as size, surface charge, loading efficiency, stability and
biodistribution can be altered for a particular application. For
example, the size of the NPs can be controlled by adjusting the
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length of the hydrophobic moiety and the length of the polymer.27
Scaling relations for this purpose have been developed.27
In addition, the surface charge, which affects particle serum
stability and cellular uptake, can be altered by controlling the
degree of substitution, the length or nature of the hydrophobic
moiety.24
Polysaccharides can be modified with a wide range of hydrophobic moieties, among them are bile acids (e.g., 5b-cholanic
acid, cholic acid and deoxycholic acid), fatty acids (e.g., palmitoyl
acid, stearic acid, oleic acid),24 cholesterol and hydrophobic drugs.
3.4
Polysaccharide–drug conjugate
The concept of polymer–drug conjugate was first introduced
in 1975 by Ringsdorf for the delivery of small hydrophobic
drugs.24 The conjugation to a polymer alters the biodistribution
and circulation time of the free drugs, which are capable after
conjugation of being selectively delivered and accumulated
at the tumor site due to the EPR effect (Fig. 3E). The benefits
of this delivery system have led to several phase I/II clinical
trials.24 This promising concept has been used for the preparation of polysaccharide–drug conjugates especially for the delivery
of non-soluble anticancer drugs as will be detailed below.
The polysaccharide–drug conjugate consists of three parts:
the water-soluble polymer, the drug and the biodegradable spacer
connecting the two. Additional components such as labeling
agents and targeting moieties may also be included. A special
attention as far as rational design of this delivery system
should be given to the spacer used. The following principles
should be considered for a successful delivery system: first, the
spacer should be stable in the bloodstream to increase circulation time but rapidly broken after cell entry.29 In addition, the
spacer should release an intact drug molecule without altering
its chemical structure.29 Furthermore, the polysaccharide itself
should be stable enough in the bloodstream.29
The final size and shape of the polysaccharide–drug conjugate greatly depend on the characteristics of the components
for example, conjugation of a hydrophobic drug results in selfassembly to a spherically shaped NP with the drug physically
trapped inside the particle.
4. Polysaccharide-based nanoparticles
4.1
Chitosan based nanoparticles
Chitosan (CS)’s nature and the ease of chemical modifications
enable multiple NP preparation schemes, among them are covalent
cross-linking, ionic cross-linking, polyelectrolyte complexation,
and self-assembly.7 Since CS is the only positively charged
polysaccharide it has been extensively used to generate polysaccharides particles as drug delivery systems.
Early work describing CS NPs for drug delivery was based
on covalently crossing CS with glutaraldehyde.7 However, since
glutaraldehyde is highly toxic, biocompatible alternatives for
covalent cross-linking have been developed such as condensation reactions with N,N-(3-dimethylaminopropyl)-N-ethyl carbodiimide (EDC),which was used to facilitate intramolecular
cross-linking of CS by natural di- and tricarboxylic acids.9
This method allows the formation of polycations, polyanions,
and polyampholyte NPs.
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Since CS is positively charged, ionic cross-linked particles
can be prepared using polyanions; among them the most widely
used is tripolyphosphate (TPP). The wide use of TPP in the
preparation of CS NPs is a result of both being non-toxic and of
the ability to modulate particle size, morphological properties
and surface charge mainly by controlling the CS to TPP weight
ratio.30 One of the first TPP cross-linked CS NPs for drug
delivery were developed by Alonso’s group based on a principle
reported previously by Bodmeier et al.7 This ionotropic gelation
technique involves the addition of an alkaline phase containing
TPP into an acidic phase containing CS.7 The formation of the
NPs upon mixing the two phases is immediate and involves
inter and intramolecular linkages created between TPP phosphates and CS amino groups.7 The characteristics of the NPs
obtained were shown to be depended on several factors such as
the concentrations of CS and TPP, CS’s purity and molecular
weight.
The group of Alonso later reported the use of these particles
for protein, oligonucleotides and plasmid DNA delivery. The
resulting CS/TPP NPs for DNA delivery were in the range of
100–300 nm depending on the Mw of the CS and showed high
physical stability and encapsulation efficiencies both for plasmid
DNA and dsDNA oligomers (20-mers), independent of CS’s
Mw.31 All obtained NPs showed high physical stability with no
detection of DNA release even after incubation with heparin as
a competitive anion. The efficiency of transfection, however,
was highly depended on the Mw of CS (Fig. 4). Unlike the NPs
based on high Mw CS, the low Mw CS/TPP particles gave high
gene expression levels in the human embryonic kidney cell line
(HEK 293 cells) 2 days post transfection, which was sustained
for up to 10 days.31 In addition, only the low Mw CS/TPP NPs
mediated a strong b-galactosidase expression in vivo after
intratracheal administration.
As discussed earlier, the positive charge of CS enables the
formation of polyelectrolyte complexes (PEC) with negatively
charged polymers such as negatively charged polysaccharides,
nucleic acids, negatively charged peptides and poly(acrylic
acid). Polyelectrolyte complexation is one of the most frequently
used methods to prepare CS NPs. PEC of CS with g-poly(glutamic acid), a natural, non-toxic, and biodegradable negatively
charged polymer, have been prepared for oral administration
of insulin.32 These NPs, which were stabilized with TPP and
magnesium sulfate, were pH sensitive and had an average size
of 218 nm in diameter. The particles were shown to be safe,
adhered to mucosal surfaces and induced a significant hypoglycemic action for at least 10 h in diabetic rats when administered orally. The bioavailability of insulin, which was determined
from plasma insulin concentration, was about 15%. The same
group also reported transdermal delivery of DNA containing CS–g-poly-(glutamic acid) NPs.9 These particles showed
improved skin penetration and enhanced gene expression in
comparison to particles solely comprised of CS and DNA.
This can be attributed to a greater density of the g-poly(glutamic acid) containing NPs, which contributed to a larger
penetration momentum into the skin barrier. Another study
describing PEC of CS and DNA was reported by Krishnendu
et al.,33 which showed that these PEC can generate immunologic protection in a murine model of peanut allergy. The mice
that received NPs containing a dominant peanut allergen gene
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Fig. 4 Ionically crosslinked CS/tripolyphosphate (CS/TPP) nanoparticles for oligonucleotide and plasmid DNA delivery. (A) and (B)
Transfection efficiency of LMW CS/TPP nanoparticles. Comparison of GFP transfection efficiency between LMW and HMW CS/TPP
nanoparticles at 10 and 20% pDNA loading, respectively, at a dose of 1 mg pDNA. (C) Macroscopic evaluation of b-galactosidase expression
in mouse lungs. Lungs were analyzed 72 h after administration of LMW CS/TPP (10% pDNA loading) and HMW CS/TPP nanoparticles
(20% pDNA loading) at a dose of 10 mg pCMV-bGal. Naked pCMV-bGal was used as control. Reprinted with permission from ref. 31. Copyright
2009 Int. J. Pharm.
produced secretory IgA and serum IgG2a and showed a
substantial reduction in allergen-induced anaphylaxis associated with reduced levels of IgE, plasma histamine and
vascular leakage. More recently, plasmid–CS PEC were applied
in the delivery of basic fibroblast growth factor (FGF-2) and
platelet-derived growth factor (PDGF-BB), which regulate cell
growth and division.34 Plasmid–CS PEC containing FGF-2/
PDGF-BB genes were injected into BALB/C mice. Several
formulations were tested, which differed in the degree of CS
deacetylation and Mw. ELISA assays performed on mice sera
showed FGF-2 and PDGF-BB expression. In addition, induction of specific antibodies against these proteins has been shown.
PEC containing highly deacetylated low Mw CS were found to
efficiently induce protein expression with minimal production of
neutralizing antibodies, which was also confirmed by histological
analyses.
Recently, PEC of ultrapure CS monomers were used for
ocular gene delivery.35 The NPs, which were created based
on a method developed by Koping-Hoggard et al.,36 had an
average size of B100 nm in diameter and a strong positive
charge. This formulation demonstrated effective transfection
of kidney cell line COS-7 in vitro and 5.4-fold higher luciferase
gene expression than polyethylenimine–DNA NPs had upon
injection to rat corneas. The preparation method of PEC of
CS and DNA was further improved by P. Artursson et al.,
who optimized the balance between stability and unpacking of
PEC containing CS oligomers of different sizes.37 PEC of CS
oligomers (or low Mw CS) and pDNA have several advantages over PEC of high Mw CS with pDNA: high Mw CS
forms extremely stable PEC with DNA, which delays the
release of DNA and therefore results in a slow onset of
action.36 In addition, these PEC are of aggregated shapes,
their viscosity at concentration used for in vivo delivery is very
high and their solubility at a physiological pH is low.36
CS drug conjugates have also been developed mainly for
the delivery of insoluble anti-cancer agents. Very recently, an
interesting pH responsive CS–drug conjugate has been introduced (Fig. 5) for photodynamic therapy.38 The system is
composed of a glycol CS backbone, a functional group
(3-diethylaminopropyl isothiocyanate, DEAP block), a photosensitizing model drug (chlorine e6 block, Ce6) and PEG.
The system is designed to form a three-dimensional supramolecular self-assembly (i.e., self-quenched state of photosensitizing drugs) at physiological pH and to destabilize into
extended random molecules (i.e., dequenched state for singletoxygen production) at lower pH as present in the tumor
surrounding due to protonation of the functional group DEAP.
The authors tested this conformational change by examining
particle sizes that have changed dramatically from 150 nm
spherically shaped NPs at pH 7.4 to 3.4 nm disentangled form
at pH 6.8. In addition, the zeta potential also changed with the
pH confirming the protonation of the DEAP group. The pHdependent change in photoactivity was confirmed by measuring
the amounts of singlet oxygen. In vitro, higher phototoxicities
of HeLa cells were observed at pH 6.8 and 6.4 than at pH 7.4.
In vivo, the CS–drug conjugate showed tumor specificity after
intravenous administration to nude mice bearing HeLa tumor
cells with a signal lasting for more than 24 hours.
Fig. 5 A pH responsive polysaccharide–drug conjugate. At physiological pH values the polysaccharide–drug conjugate self-assembles into a
nanoparticle. When reaching the acidic tumor environment, protonation of the functional group enables disassembly of the nanoparticle and
accelerated drug release.
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Chemical modification, usually by utilizing the primary
amino groups of CS, is another way to improve its physicochemical properties. For example, conjugation of hydrophobic
moieties such as deoxycholic acid and cholesterol to CS allows
solvent induced self-assembly into NPs.7 This principle was
used for preparation of 5b-cholanic acid (HGC) modified
glycol CS NPs for the delivery of the antiangiogenic peptide
RGD.39 The RGD (Arg–Gly–Asp) peptides can specifically
target avb3 integrins on angiogenic endothelial cells and therefore inhibit angiogenesis and tumor growth however, these
peptides have short half-live in vivo and thus a delivery system
is required.39 The self-assembled polymeric NPs of glycol CS
modified with hydrophobic bile acid analogs have hydrophilic
shells of glycol CS and hydrophobic cores of bile acid derivatives. The hydrophobically-modified glycol CS was prepared by
covalently attaching 5b-cholanic acid through amide formation.
The NPs had a diameter of 230 nm and loading efficiency of
>85%. The particles demonstrated prolonged and sustained
release of drug efflux and inhibition of human umbilical vein
endothelial cells (HUVEC) adhesion in vitro. In the in vivo
study, the RGD containing NPs inhibited basic (b)FGF-induced
angiogenesis and significantly decreased tumor growth and
microvessel density in comparison to the native RGD peptide.
These particles were previously used by the same group for gene
delivery and for the chemotherapeutic agents doxorubicin and
paclitaxel.9
Other chemical modifications of CS such as addition of thiol
groups and trimethylation can improve the mucoadhesive and
permeation enhancing properties of CS. Trimethylated CS (TMC)
is also characterized by increased solubility in neutral pH.
4.2
Alginate based nanoparticles
The early work reported on alginate-based NPs was focused
on the ability of alginate to form a 3D network upon ionic
inter- and intramolecular cross-linking with divalent ions.40
Since then, several preparation mechanisms have been utilized
and alginate based NPs have been developed for the delivery
of proteins, genes, anti-tubercular and anti-fungal drugs.41
The group of Gopal K. Khuller has presented several papers
describing alginate NPs for the treatment of tuberculosis, a disease
in which patient non-compliance is one of the main contributors
towards treatment failure and multi-drug resistance.42 Alginate
encapsulation of anti-tubercular drugs may present a possible
solution by improving drug bioavailability and enabling a
controlled release profile as has been reported for the antitubercular drugs isoniazid (INH), rifampicin (RIF), pyrazinamide
(PZA) and ethambutol (EMB) in comparison to free drugs.42 The
NPs were prepared by calcium induced gelification followed
by polyelectrolyte complexation with CS and were characterized by a size of about 235 nm and high encapsulation efficiencies (70–90%, 80–90%, and 88–95% for INH, RIF, and EMB,
respectively). Another advantage of the alginate NPs is the ability
to co-encapsulate multiple anti-tubercular drugs. Calcium crosslinked alginate NPs containing econazole and anti-tubercular
drugs for the treatment of murine tuberculosis were reported.42
The encapsulated drugs were detectable above minimum inhibitory concentration for 15 days after administration in lungs, liver
and spleen of the treated mice in comparison to 12–24 hours of
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the free drugs. In addition, the alginate particles managed to
reduce bacterial burden in the lungs and spleen of mice infected
with Mycobacterium tuberculosis by more than 90% at 15-fold
lower dosages in comparison to free drugs.
CS alginate NPs were used for the transmucosal delivery of
insulin.43 The NPs were prepared by ionic cross-linking of CS
and TPP, which was followed by complexation with alginate.
The obtained NPs size was affected by the overall concentration
of the added electrolyte, the mass ratio of alginate and CS, and
by the molecular weight of the alginate. Insulin was associated
to the CS–TPP–alginate NPs with loading efficiency of 41–52%.
The CS–TPP–alginate NPs were administered nasally and
exhibited a capacity to enhance systemic absorption of insulin.
The CS–TPP–alginate NPs were characterized by a longer hypoglycemic response in comparison to CS–TPP NPs. The duration
of the hypoglycemic response also depended on the Mw of
alginate.
Recently, surfactant–alginate hybrid NPs have been employed
for dual chemo- and photodynamic therapy on a murine drugresistant tumor model.44 The obtained NPs had an average size
of 70 nm measured by Dynamic Light Scattering (DLS) and
contained doxorubicin and methylene blue with encapsulation
efficiencies of 78% and 82%, respectively. Following administration to Balb/c mice bearing syngeneic JC tumors (mammary
adenocarcinoma), the dual therapy significantly inhibited tumor
growth and improved animal survival. The treatment resulted
in enhanced tumor accumulation of both doxorubicin and
methylene blue, significant inhibition of tumor cell proliferation, and increased induction of apoptosis.
Chemically modified alginate has also been developed for
the preparation of NPs. As with CS, chemical modifications
improve the physiochemical characteristics of alginate. For
example, thiolated alginate achieved by covalent attachment
of cysteine improves the mucoadhesive properties of alginate
and provides improved stability of the drug delivery system.45
Hydrophobically modified alginate derivatives have also been
produced for the preparation of self-assembled NPs for the
sustained release of vitamin D3.46
4.3 Hyaluronan based nanoparticles
HA based nanocarriers were developed using several approaches
such as HA–drug conjugates, which restore their cytotoxicity
upon cell internalization by receptor mediated endocytosis,
PEC with polycations, and ionically crosslinked NPs. Chemically modified HA has also been widely used for the delivery of
proteins, peptides and nucleotides.47 Chemical modifications
assist in prolonging HA half life however, beyond a certain level
of modification, HA loses the ability to bind its receptors.47
HA based NPs hold a major advantage over other polysaccharides based NPs, which is the ability to combine both
passive targeting by utilizing the EPR effect in tumors and
active targeting towards the HA receptors overexpressed by
the majority of tumors.
HA–drug conjugates utilize several chemical groups on HA:
the carboxylate on the glucuronic acid, the N-acetylglucosamine
hydroxyl, the reducing end and the acetyl group, which can be
enzymatically removed from the N-acetylglucosamine.13 HA–
drug conjugates are internalized via CD44 receptor-mediated
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endocytosis and the drug is released mainly by intracellular
enzymatic hydrolysis.13 In addition to targeting, HA conjugation has been used to increase drug solubility. This quality was
used for the delivery of the hydrophobic antimitotic chemotherapeutic agent, paclitaxel (PTX).13 HA–PTX conjugates were
shown to increase cellular uptake and cytotoxicity in vitro in
comparison to the free drug.13 In addition, cellular uptake of the
HA–PTX conjugate was shown to be dose dependent and CD44
specific as it could be blocked by free HA or by anti-CD44
antibodies but not by the structurally related polysaccharide,
chondroitin sulfate (see Fig. 2).13 In another study, the antitumor
activity of the HA–PTX conjugate was demonstrated in vivo.
This conjugate significantly decreased tumor burden in comparison to free PTX in mice bearing abdominal tumors of ovarian
cell line.13
Recently, HA conjugation was used for the delivery and
improved serum stability of Exendin 4.48 Exendin 4 (exenatide)
is a 39 amino acid peptide incretin mimetic that exhibits glucoregulatory activities.48 Exendin 4 has been shown to induce
glucose-dependent enhancement of insulin secretion, glucosedependent suppression of inappropriately high glucagon
secretion, slowing of gastric emptying, reduction of food intake
and body weight, and an increase in b-cell mass.48 However,
exendin 4 has a significantly short half-life, which limits its
clinical application.48 Conjugation to vinyl sulfone modified
HA resulted in 20-fold increase in serum stability in vitro
without loss of bioactivity. HA–exendin 4 conjugates lowered
glucose levels in type 2 db/db mice and the hypoglycemic effect
lasted up to 3 days after injection. In addition, insulin immunohistochemical analysis of islets in db/db mice confirmed the
improved insulinotropic activity of the HA–exendin 4 conjugates. HA conjugates have also been suggested for the treatment of acute promyelocytic leukemia (APL).49 APL patients
often relapse due to resistance to the therapy with all-trans
retinoic acid. Because of the molecular basis of APL alteration
and previous success in treating tumors with HA conjugated to
the histone deacetylase inhibitor butyric acid,49 conjugation of
HA to both all-trans retinoic acid and butyric acid has been
tested. In vitro, the HA conjugates induced growth arrest and
terminal differentiation in retinoic acid sensitive cells and
apoptosis in retinoic acid resistant cells. In vivo, HA conjugates led to a significant increase in survival time of a retinoic
acid sensitive APL murine model in comparison to that
induced by a maximum tolerated dose of retinoic acid alone.
In addition, in a retinoic acid resistant murine model, the HA
conjugates were active in contrast to retinoic acid that was
completely ineffective.49
As detailed above, PEC have been used to prepare HA
based NPs. HA–CS NPs have demonstrated the ability to
transport genes across the ocular mucosa and transfect ocular
tissue.50 The NPs were prepared by ionically cross-linking CS
with TPP, which was followed by PEC with HA. The HA-low
Mw CS particles led to high expression levels of the transfected
alkaline phosphatase in a human corneal epithelium model.
Upon topical administration to rabbits, the NPs managed to
overcome cellular barriers and were located inside the corneal
and conjunctival cells, suggesting that they penetrate the
epithelia by a transcellular pathway. In addition, the in vivo
transfection levels reached were significant.
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As with other polysaccharides, self-assembly of hydrophobicallymodified HA has been used for NP preparation. In a recent
study, HA was modified with the 5b-cholanic acid to form selfassembled NPs that combine both passive tumor targeting
based on the EPR effect and a more specific or active targeting
exploiting the affinity of HA towards CD44 (Fig. 6A).51 The
obtained NPs size could be controlled in the range of 240–430 nm
by varying the degree of substitution of the hydrophobic
moiety. In vitro, fluorescently labeled Cy5.5–HA-NPs were
detected in the cytosol of CD44 overexpressing cells (SCC7) to
a much greater extent than cells with low CD44 expression
(CV-1) (Fig. 6B). When administered systemically to tumor
bearing mice, the particles were shown to selectively accumulate in tumor sites. Smaller HA-NPs were able to reach the
tumor site more effectively than larger HA-NPs. In addition,
the concentration of the NPs in the tumor site was dramatically reduced when mice were pretreated with free HA (Fig. 6C).
This suggests an additional active targeting mechanism, beyond
the passive targeting of the EPR effect.
Another recent study demonstrated the antitumor activity
of self-assembled modified HA-NPs.52 To this end, poly(g-benzyl L-glutamate) modified HA NPs were loaded with
DOX. The NPs presented a sustained drug release pattern at
pH 5.5 and 7.4 for up to 10 days. In vitro, the NPs were tested
on both MCF-7 human breast cancer cells highly expressing
CD44 and U87 human glioblastoma cells characterized by low
CD44 expression. The accumulation of the NPs in MCF-7
and U87 was consistent with CD44 levels as demonstrated by
flow cytometry. Microscopic evaluation revealed that the NPs
were present mostly in the cytoplasm of both cell lines. In vivo
the NPs significantly suppressed tumor growth in a breast
cancer rat model, in comparison to free DOX as determined
by measuring both tumor volume and burden. In addition, the
NPs reduced the cardiotoxicity of DOX.
4.4 Dextran based nanoparticles
Several strategies have been reported for dextran based NPs,
among them are dextran–drug conjugates and self-assembly of
hydrophobically-modified dextran.5
Upon conjugation of poorly water-soluble drugs to dextran
via activation of the carboxylic groups with N,N 0 -carbonyldiimidazole, hydrophobic derivatives that self-assemble into
NPs are formed.53 The self-assembled particles were shown to
be stable at pH values ranging between 4 and 11 and with high
loading efficiency. The particle size was strongly influenced by
the degree of substitution and preparation technique.
Dextrane conjugated with hydroxyethyl methacrylate (HEMA)
was used for the preparation of NPs for siRNA delivery using
the inverse miniemulsion photopolymerization method.54 For
the preparation of the NPs, HEMA was conjugated to dextran
via a carbonate ester that is subject to hydrolysis under physiological conditions to enable biodegradation of the NPs. The
dextran–HEMA was photopolymerized with cationic methacrylate
that enables the NPs to entrap siRNAs based on electrostatic
interactions. The obtained particles were characterized by high
siRNA loading capacity, no significant cytotoxicity and biodegradability, a trait that was shown to be essential for effective
gene silencing. In addition, confocal microscopy analysis revealed
endolysosomal localization of the NPs following internalization
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Fig. 6 Self-assembled hyaluronan-based nanoparticles for active tumor targeting. (A) Structure of Cy5.5-labeled hyaluronan nanoparticles in
aqueous solution. (B) Confocal images of SCC7 tumor cells and CV-1 cells. The cells were treated with Cy5.5-labeled HA-NPs. (C) In vivo
accumulation of HA-NPs using fluorescence images. Reprinted with permission from ref. 51. Copyright 2010 Biomaterials.
into human hepatoma cell line HuH-7. In order to determine
the effect of enhanced endosomal escape on the extent of
gene silencing the authors tested photochemical internalization
and the use of influenza-derived fusogenic peptide. Photochemical internalization is a method in which amphiphilic
photosensitizers are utilized to destabilize endosomal vesicles.
Fusogenic peptides are peptides of viral origin, which are
involved in the fusion between the viral envelope and the
endosome and assist to transport the viral genome into the
cytoplasm following endocytosis.55 Both methods of endosomal escape significantly improved gene silencing.
Dextran sulfate based NPs have been mostly prepared by
PEC, exploiting the anionic nature of dextran sulfate for
electrostatic interaction with positively charged polycations
(for example, CS and polyethylenimine). Huang et al.56 prepared CS–dextran sulfate PEC for the controlled release of
vascular endothelial growth factor (VEGF). VEGF, a growth
factor, which stimulates angiogenesis and therefore desired as
a therapeutic approach for ischemic conditions, was shown to
generate new blood vessels in vivo. However, intravenously
injected VEGF was not clinically successful and implantable
controlled release devices have shown that localized and
sustained release of VEGF is required for its favorable action.
NPs of B250 nm were prepared in which the heparin binding
domain of VEGF was utilized to bind the polyanion dextran
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sulfate. The encapsulation efficiency of VEGF was high (85%)
and controlled release of active VEGF persisted for more
than 10 days. The activity of VEGF was determined by ELISA
and by the ability to stimulate endothelial cell proliferation
(mitogenic assay). PEC-containing different polycations (polyethylenimine and poly-L-lysine) were also tested, however
CS–dextran sulfate complexes were preferred because of their
biodegradability, desirable particle size, higher entrapment
efficiency, controlled release, and mitogenic activity. In a
following study by the same group57 Repifermins containing
NPs were prepared in the same manner. Repifermins is a
truncated form of fibroblast growth factor-10 that exhibits
promise in wound healing applications. The challenge of the
delivery lies in the instability of this protein. The resulted 250 nm
particles showed high encapsulation efficiency and the release
of active Repifermins was controlled for more than 10 days.
In addition, the mitogenic activity of Repifermins on human
umbilical cord vascular endothelial cells was only demonstrated for encapsulated and not free Repifermins.
Zinc stabilized complexes of dextran sulfate and polyethylenimine (PEI) have been used for the delivery of proteins,
DNA and the poorly water-soluble antifungal agent amphotericin B.58 The preparation method of these NPs was
complex coacervation, a method that was previously used for
microencapsulation.58 The effects of preparation conditions
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and composition on the physicochemical properties of the
particles have been determined.58 The sizes of the amphotericin B containing particles ranged from 100 to 600 nm, with
a zeta potential of 30 mV and drug recovery efficiency of up to
85%. Particle size was shown to be controlled by processing
parameters such as the pH of the PEI solutions, the ratio of
the two polymers and the concentrations of dextran sulfate
and zinc sulfate. The amphotericin B-containing particles
displayed no toxicity in tissue culture in contrast to free drug
and were almost as efficacious as free drug in killing Candida
albicans.
4.5
Cyclodextrin based nanoparticles
Cyclodextrin (CD) containing polymers have been used in
pharmaceutical applications since the 1980’s.15 In recent years
CD-containing polymers have been used for the formation of
NPs designed for controlled or sustained drug release, molecular absorption, tissue engineering or localized delivery of
therapeutic agents.14 Structurally, CD-containing polymers
can be classified according to location of the CD moieties in
the polymer network. The CDs can be either located in the
polymer backbone or grafted onto a preexisting polymer.14
Another classification of CD-containing polymers relies on the
obtained polymer’s charge whether it is non-ionic, cationic or
anionic.
CD-containing polycations (CDPs) have demonstrated unique
capabilities for nucleic acid delivery: they were reported to have
low in vitro and in vivo toxicities when compared with other nonCDPs such as poly-L-lysine and PEI.15
The group of Mark E. Davis had a major contribution
in the area of gene delivery using CDPs. They have created
a linear CDP and reported the formation of self-assembled
nano-sized polyelectrolyte complexes with pDNA with a cell
transfection capability compared to PEI and Lipofectamine.14
In addition, the relations between the polycation structure
and gene delivery capabilities were determined.14 Among the
structural characteristics examined were CDs size, distribution
and nature of cationic elements and the polymer size and
polydispersity.14 It was determined that low molecular weight
polymers with sufficient distance between the CD units and the
amidine cationic centers were optimal, contributing to both
high delivery and low cytotoxicity.14 Davis’s group further
improved this CD-based delivery system to create the first
clinically tested targeted NPs for siRNA delivery. This system
exploits the CD inclusion capabilities to provide the oligonucleotide polycation PEC with both steric stabilization and
targeting capabilities (Fig. 7).14
Structurally, the CDP assembles with the siRNA primarily
via electrostatic interactions. It condenses siRNA and protects
it from nuclease degradation.59 The CDs within the polymer
chains that reside on the surface of the NPs are used for assembling PEG that endows the NP with stearic stabilization by
preventing aggregation and non-specific interactions with
biological components. The PEG is conjugated to adamantine,
which forms inclusion complexes with CD. The same principle
is applied for the assembly of the targeting ligand transferrin,
which was chosen since its receptor is upregulated on cancer
cells (Fig. 7). Primary results from phase I clinical trials using
siRNA against the M2 subunit of ribonucleotide reductase R2
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Fig. 7 Cyclodextrin based nanoparticles for systemic siRNA delivery.
Schematic illustration of the delivery system. The polyethylene glycol
(PEG) molecules are terminated with adamantane (AD) that form
inclusion complexes with surface cyclodextrins to decorate the surface
of the nanoparticle with PEG for steric stabilization and PEG-TF for
targeting. Modified from ref. 59.
(RRM2) have been recently published.59 Following systemic
administration to patients with solid cancers, tumor biopsies
revealed intracellular and dose depended localization of the
NPs. In addition, specific reductions of both RRM2 mRNA
and protein levels were observed. 5 0 -RLM-RACE analyses
have shown that this reduction was mediated by an RNAi
mechanism.59 The paper describing results from the clinical
studies did not contain data regarding the effects on human
metabolism and immune response. However, following longterm administration to mice, no abnormalities in interleukin-12
and IFN-a, liver and kidney function tests, complete blood
counts, or pathology of major organs were observed.60 The
administered NPs used for delivery of EWS-FLI1 siRNAs,
demonstrated a significant and ligand specific inhibition of
tumor growth in a murine model of metastatic Ewing’s sarcoma
(Fig. 8).60 Only targeted NPs, which contained siRNA against
EWS-FLI1 (a possible transcriptional activator that plays a
significant role in the tumorigenesis of Ewing’s sarcoma), were
able to significantly inhibit tumor growth in comparison to
naked siRNA, non-targeted NPs or targeted NPs containing
control siRNA (Fig. 8A). The injection of luciferase expressing
tumor cells enabled result quantification by measuring the
tumor bioluminescent signal (Fig. 8B).
CS has been also used as a basis for the formation of CDP
for gene delivery.14 Alonso’s group has demonstrated that
pentasodium polyphosphate-mediated cross-linking of native
CS with anionic CDs results in the formation of NPs (100–200 nm
in diameter). The obtained NPs were characterized by higher
pDNA loading capacity and stabilization in comparison to
NPs produced without CDs and presented enhanced cellular
uptake and protein expression.14
Other CD-based NPs have been developed for the delivery
of drug and nucleic acids, among which are CD-based polyrotaxane–pDNA NPs and CD-based dendrimers.14
4.6 Arabinogalactan based nanoparticles
Arabinogalactan has been used as a carrier for drug conjugates.
Recently, arabinogalactan–folic acid–drug conjugates for targeted
delivery and activated drug release were prepared.17 The targeted
nanovehicle was formed by conjugation of folic acid and the
anticancer drug methotrexate to arabinogalactan. The use of
folic acid as a targeting ligand derives from the fact that cancer
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Fig. 8 Tumor growth inhibition in a murine model of metastatic Ewing’s sarcoma as a result of sequence specific knockdown of EWS-FLI1 by a
targeted cyclodextrin based siRNA delivery. (A) Bioluminescence imaging of NOD/scid mice treated twice weekly with formulated siRNA for
4 weeks. Starting immediately after injection of TC71-LUC cells, mice were treated with formulations containing siRNA targeting EWS-FLI1
(siEFBP2) or a nontargeting control sequence (siCON1). (B) Growth curves for engrafted tumors. The median integrated tumor bioluminescent
signal (photons s1) for each treatment group (n = 8–10) is plotted versus time after cell injection. Treatment groups: A, 5% (w/v) glucose only
(D5W); B, naked siEFBP2; C, targeted, formulated siCON1; D, targeted, formulated siEFBP2; and E, nontargeted, formulated siEFBP2.
Reprinted with permission from ref. 60. Copyright 2005 Cancer Res.
cells overexpress receptors for nutrients in order to maintain
their fast-growing metabolism.2 One of these receptors is
the folate receptor, which is overexpressed in malignant cells
including ovary, brain, kidney, breast, colon, and lung.17
Another advantage of using nutrient receptors as receptor
targets is that they enable internalization of the nanocarrier
via receptor-mediated endocytosis. The activated drug release
was achieved by linking methotrexate to arabinogalactan by
an endosomally cleavable peptide Gly-Phe-Leu-Gly (GFLG).
The nanocarrier displayed significant cytotoxic activity to
folate receptor overexpressing cells in comparison to folate
receptor deficient cells.
4.7
Pullulan based nanoparticles
Pullulan has been shown to self-assemble into NPs after modification by hydrophobic molecules such as cholesterol and stearic
acid. In addition, pullulan–drug conjugates and covalently
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crosslinked pullulan NPs have been reported. Overall, pullulanbased NPs have been used for the delivery of proteins, anticancer
drugs, imaging agents, and nucleotides.
Akiyoshi K. et al. have been developing self-assembled NPs
of cholesterol-modified pullulan for more than a decade. The
hydrophobically modified pullulan self-assembles in water into
monodisperse and colloidally stable NPs.61 Akiyoshi and
coworkers examined the effect of various modified cholesterol
moieties, different Mws of pullulan and different degrees of
substitution of the cholesterol moiety.9 The self-assembled
particles demonstrated the ability to complex various substances including soluble proteins such as insulin by mainly
hydrophobic interactions.61,62 The complex between the NP and
protein solution was easily formed by simply mixing the two
components.62 The particle size did not change after complexation with the proteins. The complexation with insulin occurred
faster than with larger proteins such as a-chymotrypsin or BSA.
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These particles showed high colloidal stability with no dissociation of insulin or precipitation. The complexation greatly contributed to the thermal stability of insulin (even after heating for
6 h at 90 1C) and protected insulin from enzymatic degradation.
In addition, in vivo experiments demonstrated preservation of
insulin’s bioavailability.62
The ability of these NPs to thermally stabilize proteins
was further investigated. The molecular chaperon-like activity
was demonstrated on proteins using a system consisting of
the cholesterol–pullulan NPs and b-CD. Capture of heatdenatured unfolded protein and the release of the refolded
form were achieved. In addition, an irreversible protein aggregation upon heating was completely prevented, recovering
almost 100% of protein activity.9
Intracellular protein delivery by self-assembled NPs of cationic
cholesteryl group-bearing pullulans was also demonstrated.63 In
this study, the cationic derivative of the cholesterol was used
instead of cholesterol since it better facilitates cellular uptake.
While particle size did not change significantly upon replacing the
cholesterol with the cationic cholesteryl group, the zeta potential
became positive (+7.7 0.1 mV in comparison to 1.3 0.4).
Additionally, the binding constant to the model protein, BSA,
significantly grew. The cholesteryl-pullulan NPs demonstrated a
more effective internalization of proteins into HeLa cells in
comparison to cationic liposomes and a protein transduction
domain (PTD)-based carrier. Upon cell internalization, the
protein containing NP dissociated and the protein was released.
The NPs were also tested in the area of cancer vaccine
development. Clinical studies demonstrated that subcutaneous
injection of NPs carrying the cancer antigen HER2 (CHP–
HER2) or NY-ESO-1 (CHP–NY-ESO-1) induced an antigenspecific CD8+ cytotoxic T lymphocyte response and antibody
production.61
Recently, the cholesteryl–pullulan NPs were used as an antigenic
protein delivery system for adjuvant free intranasal vaccines.61
Intranasal delivery of a non-toxic subunit fragment of Clostridium
botulinum type-A neurotoxin using these NPs resulted in antigen
adherence to nasal epithelium. The NPs internalized into the nasal
epithelium immediately after the intranasal administration and the
antigen was later dissociated from the NPs in a controlled manner
in the nasal epithelial cells. In addition, the chaperon-like activity
reported for other proteins was demonstrated in this system as well.
In vitro, circular dichroism analysis showed that the secondary
structure of the antigen was changed after the incorporation into
the NPs, but recovered after it is released. In order to prove
induction of antigen specific respiratory immune response,
flow cytometric and immunohistochemical analyses were utilized
and demonstrated that within 6 h after NPs administration, the
antigen released from the nasal epithelium was taken up by
CD11c+, dendritic cells. The NPs also induced vigorous
botulinum-neurotoxin-A-neutralizing serum IgG and secretory
IgA antibody responses. Furthermore, intranasal immunization
of tetanus toxoid with the NPs induced strong tetanus-toxoidspecific systemic and mucosal immune responses.
5. Polysaccharide coated nanoparticles
Surface modification of NPs with polysaccharides has remarkable advantages for drug delivery systems: it endows the NP
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with long-term circulation and increased stability. The prolonged circulation time can be attributed to reduction in
protein adsorption and opsonisation. In addition, it provides
targeting abilities: pullulan coating has been used for hepatic
delivery, alginate and CS coating have been used for mucosal
delivery, mannan coating facilitated uptake by macrophages
and HA coated NPs have been used to target CD44- and
CD168-overexpressing cancer cells. Another advantage is the
‘build-in’ cryoprotection feature that has been reported for
several polysaccharides. Furthermore, polysaccharide coated
liposomes demonstrated reduced permeability to water soluble
encapsulated materials and protection from degradation by
lipases.64
Coating NPs with polysaccharides can be achieved by adsorption, incorporation, copolymerization or covalent grafting65 and
has been reported for many polysaccharides, among which are
pectin, pullulan, mannan, HA, heparin, CS and dextran.65
5.1 Chitosan coated nanoparticles
The cationic nature of CS enables its adherence to mucosal
surfaces. In addition, the ability of CS to open tight junctions
between epithelial cells has also been demonstrated.7 These
characteristics make CS appealing as a coating agent of nanocarriers designed for mucosal delivery. For example, CS coated
poly-e-caprolactone (PECL) NPs allowed the bioavailability
of the anti-inflammatory drug, indomethacin, in the cornea
and aqueous humor following topical ocular instillation.65 In
addition, CS coating of liposomes (chitosomes) enhanced mucosal
adhesion in rat intestine following oral administration.7 Chitosome
formulations used for the oral delivery of insulin and calcitonin (in separate studies) induced significantly more substantial and prolonged decreases in blood glucose and calcium
levels, respectively, relative to uncoated liposomes.7
CS coating can be used to replace the cationic polymers and
lipids currently used for nucleic acid delivery thus overcoming
the toxicity, which is the major obstacle in using these compounds. CS coated NPs can interact with nucleic acids, improving
the particle loading efficiency and transfection properties.65
Structural benefits of coating NPs with CS have also been
demonstrated: CS coated liposomes were more stable in simulated
gastric fluids in comparison to uncoated liposomes.7 CS coated
PECL particles also demonstrated enhanced physical stability.7
In addition, the CS coating also facilitated the redispersion of
lyophilized PECL nanoparticles.7
5.2 Hyaluronan coated nanoparticles
The HA capsule of group A streptococci enables it to escape
the host immune response66 and also provides long-term circulation. This feature has been successfully adopted for the delivery
of mitomycin C (MMC) using HA-coated liposomes (tHA-LIP)
(Fig. 9A).67 tHA-LIP were 7- and 70-fold longer circulating in
comparison to uncoated liposomes and free MMC, respectively,
in 3 murine tumor models: BALB/c bearing C-26 solid tumors,
C57BL/6 bearing B16F10.9 and D122 lung metastasis.67 The HA
on the tHA-LIP was covalently attached using EDC via the
glucuronic carboxylate to phosphatidylethanolamine in the preformed liposomes. The tHA-LIP demonstrated slower drug
efflux and higher encapsulation efficiency. In addition, since the
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Fig. 9 Hyaluronan coated nanoparticles. (A) A schematic illustration of HA-coated liposome. (B) PTX-GAGs. (C) LbL NP with HA as an outer
layer. Adapted from ref. 25.
effect of tHA-LIP is CD44 dependent, these NPs demonstrated
significantly higher cytotoxicity in vitro on CD44 overexpressing
cells and increased drug accumulation in tumors in vivo. The
latter resulted in decreased metastasis, inhibition of tumor
growth and prolonged survival. These effects were later demonstrated with doxorubicin loaded tHA-LIP.3 This study also
compared the tHA-LIP to poly(ethyleneglycol) (PEG) coated
liposomes (‘‘stealth’’ liposomes). This was done since PEGylation,
a common hydrophilic surface modification of drug delivery
systems, demonstrated prevention of recognition by the immune
system. The tHA-LIP were long circulating more than all tested
controls including uncoated and PEGylated liposomes in healthy
and tumor bearing mice. Furthermore, the HA coating managed
to prolong the circulation time without activating the complement, which is associated with PEG coated NPs.68
Recently, HA coated lipids–paclitaxel clusters (PTX-GAGs)
and MMC-GAGs have been prepared for selective tumor
targeting (Fig. 9B).69,70 The aqueous insolubility of PTX was
utilized in this preparation by mixing it with lipids that selfassembled into nano-sized clusters. The clusters were then
covalently coated via EDC with HA to facilitate targeting of
CD44. When tested in vivo, these cluster particles induced
tumor arrest in a murine model of colon adenocarcinoma and
were significantly more potent than free PTX and Abraxanes,
a commercially available PTX formulated in NPs. Similarly,
head and neck tumors expressing CD44 were targeted using
MMC-GAGs and showed superior therapeutic outcomes
compared to free MMC.70
Targeting of CD44 was also demonstrated for liposomes
decorated with HA oligomers.71 Unlike the previously described
tHA-LIP preparation, the HA oligosaccharides were conjugated
by reductive amination to phosphatidylethanolamine prior to
liposome preparation. The oligosaccharide decorated liposomes
demonstrated CD44 dependent uptake that could be blocked by
both free HA and anti-CD44 antibodies. Liposome uptake
depended on ligand density however, as little as 0.1 mol% HA
managed to facilitate targeting. In addition, doxorubicin
encapsulated in the oligosaccharide decorated-liposomes was
significantly more cytotoxic to CD44 overexpressing cells in
comparison to the free drug.
Another clinically relevant advantage of the HA coated
NPs is cryoprotection.72 Cryoprotection provides the liposomes with a longer shelf life since it prevents the reversion
of lyophilized unilamellar liposomes to multilamellar liposomes upon rehydration. The tested lyophilized tHA-LIP
demonstrated the ability to retain the same dimensions, zeta
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potential, encapsulation efficiencies and half-life of drug release
of the original systems upon redispersion. HA cryopreservation
is possibly by providing substitute structure-stabilizing H-bonds.
In a recent report, electrostatically assembled LbL NPs for
cancer applications with an outer layer of polysaccharides
have been presented (Fig. 9C).25 The authors tested the effect
of NPs stabilization on the biodistribution following systemic
delivery. The LbL NPs were built on a core template of gold
NPs (AuNPs) or quantum dots (QD) and were composed of
dextran sulfate (DXS) and poly-L-lysine (PLL) layers with an
outer layer of DXS or HA. In vivo, NPs showed increased
stability when larger number of layers was used. The most
outer layer was shown to be of significant importance to both
biodistribution and non-specific uptake. An outer layer of HA
resulted in a prolonged circulation time and low accumulation
in the liver. The authors tested the NPs and nanofilms in vivo
stability by administering NPs with a QD705 core that can be
tracked systemically, together with the incorporation of a layer
of labeled low Mw PLL. The authors assumed that in the case
of in vivo instability of the NPs, the incorporated low Mw PLL
would be filtered by the kidneys and detected in the bladder
shortly after administration. Hence, early detection of high
concentrations of the incorporated labeled PLL would indicate
in vivo instability.
The passive tumor targeting ability via the EPR effect was
tested post systemic intravenous injection in nude mice bearing
subcutaneously induced tumors (Fig. 10A). 24 hours post
injection, the HA coated NPs were detected in the tumors
(Fig. 10B) and the time dependent NP signal from the tumors
was shown (Fig. 10C).
5.3 Heparin and dextran coated nanoparticles
Heparin has been shown to inhibit complement activation at
different stages by increasing the activity of protein H.73 Since
the activation of the complement system plays an important
role in opsonization and uptake of particles by the mononuclear phagocyte system (MPS), surface modification of NPs
with heparin seemed promising.73 Indeed, coating NPs with
heparin significantly promoted long-term circulation of these
carriers.73 The heparin modified poly(methyl methacrylate)
(PMMA) NPs significantly elongated the in vivo half life.73 While
the circulation time of the uncoated poly (methyl methacrylate)
particles was only 3 minutes, the heparin-coated particles circulated for more than 48 h post an initial phase of elimination from
the blood with a half-life of 5 h. The heparin coated NPs were
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Fig. 10 Blood circulation and tumor targeting of optimized LbL nanoparticles. (A) Blood circulation profiles of QD705, QD705/PLL/(DXS/
PLL)3/HA, and QD705/PLL/(DXS/PLL)3/DXS. The longer persistence of QD705/PLL/(DXS/PLL)3/HA in the bloodstream corroborates their
superior stability and biodistribution profile. (B) Enhanced permeation and retention (EPR)-based targeting of solid KB tumors induced
subcutaneously on both hind flanks using QD705/PLL/(DXS/PLL)3/HA. Image is taken at 24 h time point. (C) Time-dependent accumulation of
QD705/PLL/(DXS/PLL)3/HA in KB tumors. Accumulation of the nanoparticles in tumors is transient and typical of EPR dominated targeting.
Reprinted with permission from ref. 25. Copyright 2011 Nano Lett.
still detectable in the plasma at 72 h. In addition, the dextrancoated NPs were also eliminated very slowly over 48 h. In vitro,
the heparin or dextran coated particles were also demonstrated
to be slower in the uptake by a macrophagic cell line in
comparison to uncoated particles.73 The steric barrier formed
by dense brush-like arrangement of the attached polysaccharide
chains could contribute to the long-circulating properties of the
heparin (or dextran) coated PMMA NPs.
Recently, an artificial oxygen carrier based on a polysaccharide
decorated NP was demonstrated.74 The core–shell NPs, developed as red blood cell substitutes, were covered with a long
brush of polysaccharides (heparin, dextran or dextran sulfate)
and demonstrated very low complement activation. The NPs
were obtained by using a redox radical polymerization mechanism in aqueous medium, which was followed by adsorption
or coupling of hemoglobin. In addition, the anticoagulant
properties of heparin were preserved upon coating the NPs
with heparin. When benzene tetracarboxylic acid (BTCA) was
used as a coupling agent for hemoglobin to dextran-coated
NPs, the loading capacity showed a 9.3-fold increase over NPs
in which BTCA was not used. The modification of NPs by
BTCA slightly increased complement activation; however, this
activation was reverted by the further addition of hemoglobin.
The bound hemoglobin preserved its ability for exchanging
oxygen.74
6. Summary and discussion
The variety of naturally occurring polysaccharide properties
has been successfully utilized to create multiple nano-sized
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Chem. Soc. Rev., 2012, 41, 2623–2640
nanomedicines. The advantages of polysaccharides enable the
preparation of nanocarriers for the delivery of proteins, peptides,
antibiotics and nucleic acids using several administration routes.
In addition, preliminary results from the first phase I clinical trial,
using polysaccharide-based nanocarriers, have been presented.
Polysaccharides have been shown to posses several favorable traits in comparison to synthetic polymers currently used
for drug delivery. Unlike synthetic polymers that can accumulate in the body above a certain Mw that enables renal
clearance, polysaccharides have known biodegradation routes
in which many of the specific involved enzymes have been
identified.8,12 In addition, for nucleic acid delivery purposes in
which a positive charge is required to facilitate electrostatic
interactions with the negatively charged nucleic acids, both CS
and cyclodextrin containing polymers have been shown to be
significantly less toxic than the currently used cationic polymers
and lipids.15 As far as prolonging circulation time, polysaccharides
have advantages over the currently used PEG not only related to
biodegradation, for example, HA coating of liposomes manage
to elongate the circulation time without the adverse complement
activation observed for PEGylated liposomes.68
Naturally, there are also disadvantages of using polysaccharides for drug delivery purposes, as polysaccharides
are natural materials the final product may not be consistent.
This may lead to batch variations, an unwelcome feature when
aiming for pharmaceuticals. Nevertheless, several reports for
polysaccharide production of defined size and content using
in vitro recombination techniques are available.75 In addition,
the natural nature of polysaccharides can facilitate specific and
unwanted reactions with the immune system. For example, it
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has been reported that low Mw HA fragments can induce
immune stimulation by interactions with Toll like receptors;
however, so far no immune stimulation was reported when
using NPs containing low Mw HA.68
The large variety of polysaccharides especially when also
considering all possible modifications gives the advantage of
obtaining defined materials for many specific applications but
it can also be confusing when trying to choose the appropriate
polysaccharide for a specific task. To this end, several issues
should be taken into consideration such as the requested drug
characteristics, the route of administration, the need for specific
targeting and more. For example, several polysaccharide
coatings have been shown to prolong the circulation time
(i.e. hyaluronan, heparin and dextran), which specifies them
for systemic applications. Polysaccharides chosen for colon
delivery should be mucoadhesive, stable in the rough environment of the stomach and biodegradable by the specific enzymes
found in the colon (i.e. dextran, amylose and pectin).4 Upon
aiming for specific targeting there are also several options as
pullulan can be used to target hepatocytes; alginate and CS are
mucoadhesive and can target mucosal surfaces such as vagina,
rectal and lungs; mannan can facilitate macrophage targeting,
and HA is the primary ligand for CD44 and CD168, and can
target many types of cancer cells, stem cells and lymphocytes.
Further understanding of the characteristics of polysaccharides
and the mechanisms involved in drug delivery will result in
improved drug delivery systems tailor-made for a particular
application.
Acknowledgements
The authors wish to thank Dr Micha Fridman and Dr Michael
Gozin from the Tel Aviv University School of Chemistry for
their help in the presentation of the polysaccharide chemical
structures.
S.M. thanks Tel Aviv University Nano center for PhD fellowship. This work was supported in part by grants from the
Alon Foundation, the Marie Curie IRG-FP7 of the European
Union, Levy Family Trust and Israel Science Foundation
(Award #181/10) to DP.
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